Magnetic Resonance Imaging for medical and diagnostic applications is normally carried out at magnetic fields of at least 1 Tesla (10,000 gauss), which corresponds to an RF frequency of 42 MHz for the nuclear magnetic resonance (NMR) signal of protons in water which is the major constituent of the body. The high field results in both better polarisation of the hydrogen nuclei and a strong signal at the higher detection frequency. This results in good contrast and spatial resolution. Whole body imaging at this field requires a large solenoid magnet, usually superconducting with high homogeneity typically a few parts per million over a 40 cm sphere. The patient/subject is placed on a bed which slides into the bore of the magnet. This can be a claustrophobic and unnerving experience. It is made worse due to the noisy operation of the gradient coils as images are acquired over a period of typically 15 minutes. As a result, as many as 20-40% of patients decline the procedure.
Alternative methods use so-called open geometry electromagnets operating at lower fields, typically 2,000-3,000 gauss. The reduction in field and operating frequency gives lower strength and noisier NMR signals for the detection system to acquire, with a consequent loss of image quality.
The limit in using lower fields is set by the ability of the electronic amplifiers and the resonant detection circuits to detect the signals with an adequate signal to noise ratio.
Another approach to detecting the NMR signals for MRI at low fields, developed in recent years but not yet used in practical systems, is to use a Superconducting Quantum Interference Device known as a SQUID. SQUIDs usually operate at very low fields, 1-200 gauss being typical, and have the twin advantages of very low inherent device noise and superconducting input circuits which have, under the right conditions, no Johnson or thermal noise. Squids are ideally suited to acquiring the very small NMR signals at low magnetic fields and thus allow the use of open geometry magnet systems, which have the potential to offer the patient a more friendly environment during the procedure as well as other benefits described below.
In an NMR or an MRI instrument, the sample or the subject is placed in a magnetic field which results in a polarisation of hydrogen nuclei, with the result that more nuclei point in the direction of the field rather than against it. This polarisation is quite weak for protons in water. It corresponds to parts per million at room temperature at 1 Tesla. To perform NMR, a short radio frequency (RF) pulse at the resonant frequency tips the polarisation to 90° from the field about which the polarisation rotates for a short period at the resonant frequency. It is then detected with suitable RF amplifier electronics. The polarisation decays and reverts back to point in the original direction with a characteristic time constant known as T1. During this period, the energy of the tipped magnetisation is given up to the local environment. That is to say, the nuclear spins are exchanging energy with that local environment. The speed with which energy is exchanged is a measure of the interaction of the nuclei with the local environment.
There is a second way in which the signal decays, which depends on the interaction between the nuclear spins. The so-called spin-spin interaction, in which the spin polarisation becomes spread out. If some atoms/nuclei are in different magnetic fields compared to others, their rotation frequency is slightly different. As the nuclei in different fields rotate at different speeds, the net rotating magnetisation is reduced as different moments start to point in different directions. There is a reducing net magnetic moment. This time constant for interaction between spins is known as T2 if the loss of magnetic moment is due to spin spin interaction or T2* if it is due to poor external magnetic field homogeneity.
It has been found that there are differences in T1 between tissue types and tissue conditions. In particular, it has been discovered that it is possible to detect cancerous tissue by the difference in its T1time constant. These T1differences are most pronounced at low magnetic fields well below 100 Gauss.
In conventional nuclear magnetic resonance imaging a strong and highly uniform magnetic field of typically 1 Tesla (10,000 Gauss) or more is used to magnetise the subject and a series of RF pulses is applied to polarise the spins and detect the signal from these. The frequency at which this signal is detected is directly proportional to the local magnetic field. By applying magnetic field gradients to the subject, the signal frequency becomes dependant on position, with spins in higher fields precessing at higher frequencies to those in lower fields. By applying a set of field gradients in all three dimensions, a three-dimensional picture of the subject can be acquired. In practice, gradients across one plane are used to select a planar slice through the subject. This is then analysed by gradients of field in the plane of the slice to produce a two-dimensional picture of the subject in that plane.
The magnitude of the gradient must be sufficient to disperse the spin frequencies by a sufficient amount to allow data to be acquired in the short time available, but not so large as to disperse and depress the detected signal. The background field must be uniform so that variations in the background field do not impinge on the image quality. Typically, figures might be for a 1. Tesla magnet field a gradient of 1000 ppm or 1 Milli T across the subject, requiring a background homogeneity of 2-5 ppm. Building large magnets to accommodate a whole person with a field homogeneity of a few ppm over say a 400 mm sphere is an expensive and difficult task. However, for a system running at lower field, the same field gradients for imaging are required. That is 1 about milli T across the image area, since the same frequency changes are required. This means that the homogeneity of a magnet running at 0.02 Tesla need only be 100 ppm, a figure which is much easier to achieve during magnet construction.
Two methods of using SQUIDs to perform both NMR and MRI have been proposed and demonstrated at low fields. In the first approach, known in the art, for example in the published work of John Clarke of Berkley University USA, the SQUID is used to detect the magnetic moment directly using a DC coupled input coil. The field is typically 0.1 milli Tesla (1 gauss) and frequency 4.2 KHz. In order to get sufficient signal strength it is usual to apply a pre-polarising DC field pulse of about 3000 gauss for a period which must be about T1 or longer to magnetise the subject after which the NMR/MRI pulse sequence and data acquisition can take place.
This procedure has the advantage that the differences in T1 between tissue types and conditions are most favourable at fields of 10 gauss or below. This is illustrated in FIG. 1 which shows the expansion of T1 times at the lowest fields.
The disadvantages of this approach is first the pre-polarising pulse is difficult to apply and secondly the technique is vulnerable to weak DC magnetic field disturbances which distort the image, as well as a susceptibility to RF and AF interference due to the direct DC coupling of the input coil to the SQUID. The technique would therefore normally have to be performed in a shielded room environment.
An alternative technique is to use a resonant superconducting input circuit which is AC coupled to the SQUID input. A frequency of 600 to 800 Kilo Hz and a field of 150 to 200 gauss is used in this procedure. No pre-polarising field is required and the use of a tuned AC coupled input greatly reduces the susceptibility to unwanted interference. This approach has been pioneered by Dr Hugh Seton at Aberdeen University.
However, at this field the T1 times are shorter and less differences in T1are observed between tissue types.